This invention relates imaging by use or vibratory energy such as acoustical energy and, in particular, to production of color flow images based on the Doppler shift of ultrasonic signals.
There are a number of modes in which vibratory energy, such as ultrasound can be used to produce images of objects. The ultrasound transmitter may be placed on one side of the object and the sound transmitted through the object to the ultrasound receiver placed on the other side ("transmission mode"). With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of amplitude of the ultrasound that reaches the receiver ("attenuation" mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver ("time-of-flight" or "speed of sound" mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which brightness of each pixel is a function of amplitude of the ultrasound reflected from the object back to the receiver ("reflection" "backscatter" or "echo" mode). The present invention relates to a backscatter method for producing ultrasound images.
There are a number of well known backscatter methods for acquiring ultrasound data. In the so-called "A-scan" method, an ultrasound pulse is directed into the object by the transducer and the amplitude of the reflected sound is recorded over a period of time. The echo signal amplitude is proportional to the scattering strength of the reflectors in the object and the time delay is proportional to the range of the reflectors from the transducer. In the so-called "B-scan" method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-scan method and their amplitude is used to modulate the brightness of pixels on a display at the time delay. With the B-scan method, enough data are acquired from which an image of the reflectors can be reconstructed.
In the so-called C-scan method, the transducer is scanned across a plane above the object and only the echoes reflecting from the focal depth of the transducer are recorded. The sweep of the electron beam of a CRT display is synchronized to the scanning of the transducer so that the x and y coordinates of the transducer correspond to the x and y coordinates of the image.
Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements sandwiched between a pair of electrodes. Such piezoelectric elements are typically constructed of lead zirconate titanate (PZT), polyvinylidene difluoride (PVDF), or PZT ceramic/polymer composite. The electrodes are connected to a voltage source, and when a voltage waveform is applied, the piezoelectric elements change in size at a frequency corresponding to that of the applied voltage. When a voltage waveform is applied, the piezoelectric element emits an ultrasonic wave into the media to which it is coupled. Conversely, when an ultrasonic wave strikes the piezoelectric element, the element produces a corresponding voltage across its electrodes. Typically, the front of the element is covered with an acoustic matching layer that improves the coupling with the media in which the ultrasonic waves propagate. In addition, a backing material is coupled to the rear of the piezoelectric element to absorb ultrasonic waves that emerge from the back side of the element so that they do not interfere. A number of such ultrasonic transducer constructions are disclosed in U.S. Pat. Nos. 4,217,684; 4,425,525; 4,441,503; 4,470,305 and 4,569,231, all of which are assigned to the instant assignee.
When used for ultrasound imaging, the transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages (apodizing). By controlling the time delay (or phase) and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements combine to produce a net ultrasonic wave that travels along a preferred beam direction and is focused at a selected point along the beam. By controlling the time delay and amplitude of the applied voltages, the beam with its focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (receiver mode). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the signal from each transducer array element.
This form of ultrasonic imaging is referred to as "phased array sector scanning", or "PASS". Such a scan is comprised of a series of measurements in which the steered ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, the transmission and reception are steered in the same direction (.theta.) during each measurement to acquire data from a series of points along an acoustic beam or scan line. The receiver is dynamically focused at a succession of ranges (R) along the scan line as the reflected ultrasonic waves are received. The time required to conduct the entire scan is a function of the time required to take each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio. For example, a total of 128 scan lines may be acquired over a 90 degree sector, with each scan line being steered in increments of 0.70.degree.. A number of such ultrasonic imaging systems are disclosed in U.S. Pat. Nos. 4,155,258; 4,155,260; 4,154,113; 4,155,259; 4,180,790; 4,470,303; 4,662,223; 4,669,314 and 4,809,184, all of which are assigned to the instant assignee.
Techniques for measuring blood flow in the heart and vessels using the Doppler effect are well known. Whereas the amplitude of the reflected waves is employed to produce black and white anatomical images of the tissues, the frequency shift of the reflected waves may be used to measure the velocity of reflecting scatterers from tissue or blood. Color flow images are produced by superimposing a color image of the velocity of moving material, such as blood, onto the black and white anatomical image. The measured velocity of flow at each pixel determines its color.
A major difficulty in obtaining Doppler effect measurements of reflected ultrasonic waves from blood is that the received echo signal typically contains a large component produced by stationary or slowly moving tissues, whereas blood reflects ultrasound very weakly. The stationary tissues do not produce any frequency shift in the reflected waves and these components can be easily filtered out without affecting the flow measurement. However, the reflections produced by the moving tissue due to cardiac or respiratory motion are frequency shifted and may completely overwhelm signals from slowly flowing blood. Prior systems provide a "wall filter" which is manually adjusted by the operator to filter out a narrow band of frequencies in the echo signal centered on the carrier frequency where static signals lie.
If low velocity flow is to be detected, a very narrow wall filter must be used so as not to eliminate the blood flow signal. In general, a narrow filter implemented digitally requires more filter coefficients than a broad filter, and since the number of echo signal samples available at the output of the filter is the number of input samples minus the number of filter coefficients, plus 1, this means that when narrow wall filters are employed fewer echo signal samples are available to calculate flow. For example, if a signal with 16 digital samples is filtered using an 8-coefficient FIR filter, only nine samples are available at the filter output from which the flow may be estimated. This is undesirable because error in the flow estimate due to noise is reduced as the number of data samples increases. Thus, as the wall filter is narrowed to increase the signal from low velocity flow, measurement accuracy is reduced.